Coded aperture masks consist of a pattern of apertures or pin holes in a material that has a high attenuation coefficient for a type of radiation used in diagnostic nuclear medicine imaging. Where, for example, gamma-rays, the array of apertures is arranged in a material such as tungsten which, typically, is 1-2 mm in thickness. The mask is formed by about 88 000 apertures arranged in a pre-determined manner on and extending through the tungsten sheet.
Coded aperture masks may be and are used as an alternative to various types of collimators, particularly lead collimators, in gamma-ray imaging. Lead collimators are, essentially, grid-like screens made of lead. The apertures of the grid are configured to permit transmission of parallel or near parallel gamma-rays produced by a gamma radiation source to a detector or imaging means which is typically a gamma camera. Lead and other collimators generally suffer from low resolution and attempts to increase resolution result in lowered efficiency. It is for this reason that attempts are being made to use coded aperture masks to replace lead collimators.
In addition, coded aperture masks have the potential to increase the signal-to-noise ratio (SNR) of the imaging system [1], and can, thus, theoretically be applied advantageously to diagnostic imaging in nuclear medicine. The increased SNR can be manipulated to improve image resolution, to shorten imaging time, or to reduce the patient's dose of radioactivity. The advantages of this are self evident.
Coded aperture masks have been used extensively in astrophysics, where far-field imaging conditions apply. Such conditions allow for the acquisition of images that are close to perfect for two-dimensional (2D) noise-free data [2]. Unfortunately the same cannot be said for the near-field conditions of nuclear medicine where corruption of the image by near-field artifacts is a universal problem.
Previous research has provided an indication of characteristics of apertures that are optimal for the purposes of nuclear medicine [3]. A reduction of near-field artifacts can be achieved by taking a second image with a rotated aperture, and by then summing the two sets of data obtained [4]. The use of an array of limited field-of-view coded apertures has also been shown to have the potential to significantly reduce near-field artifacts [5].
Coded aperture imaging requires that for each point of the source, the aperture pattern must be projected onto a detector. This results in overlapping aperture patterns, each shifted and weighted according to the location and the intensity of the specific point source that projected the pattern [6]. Theoretically, this acquisition process is modeled by convolving the source with the aperture pattern. The image is reconstructed by correlating the encoded data with the original coded aperture pattern [6]. This pattern is designed such that a unique reconstruction exists.
Convolution implies that a point source must be imaged equally by each pinhole of the coded aperture, without a change in intensity, and with the image of the point source falling directly below the pinhole. The decoding procedure performs correctly under these conditions, but in practice the convolution model does not hold. The near-field conditions of nuclear medicine introduce artifacts to the image. Further, the thickness of the coded aperture contributes to near-field artifacts, by collimating gamma-rays that have high angles of incidence.
With respect to image resolution, the pixel size is typically related to the size of the projection of the smallest hole in the coded aperture. The size of the smallest hole is typically designed in relation to the resolution of the gamma camera. This means that a gamma camera with 10× the resolution of existing gamma cameras, for example, can theoretically have a coded aperture that matches the 10× improvement. However, due to collimation artifacts, the minimum size of the hole is limited by the thickness of the aperture material. Similarly, the thickness of the coded aperture also constrains the manufacturing technique that can be used. Typically in laser drilling, the dimensions of the holes have to be greater than the thickness of the material, whilst a thickness of 1 mm, for example, is generally unsuitable either for etching or for deposition.
More importantly, the resolution of the resultant image is constrained, inter-alia, by the dimensions of the holes, which are in turn constrained by the thickness of the aperture material which is related to the attenuation properties of that material. If a gamma-ray passes through opaque aperture material of density ρ, with an attenuation coefficient μ specific for a given element at the energy of interest, and an effective thickness rm, the transmission t of the aperture material is given by:t=e−ρμrm  (1)
The ability of the aperture material to block gamma-rays is then given by the attenuation α:α=1−t  (2)
For a given source of radioactivity, with an associated energy, the thickness of the coded aperture material is typically chosen to give an attenuation of more than 90%, frequently 99%.
In addition to the above and with respect to image resolution, it is necessary to consider a point source that is projected through an infinitely small pinhole onto a perfect detector. If the projection is recorded by a single pixel of the detector, the representation will be correct. If the projection falls on a boundary between neighbouring pixels, counts of radioactivity will be distributed equally between those pixels. The total number of counts remains unchanged, but the measured peak is no longer representative of reality.
This problem is known as the ‘partial volume effect’ [10], and is related to the digitisation of an analogue signal. A solution is to increase the radius of the pinhole, such that the projection of the point source illuminates an area that corresponds to at least 2×2 pixels of the detector [6]. In this manner, one pixel is always fully illuminated, and the measured peak will be correct.
With respect to image resolution, the pixel size is typically related to the size of the projection of the smallest hole in the coded aperture. The size of the projection of the smallest hole is typically designed in relation to the resolution of the gamma camera, with the projection typically occupying the same area as at least a 2×2 array of detector pixels [6], in order to counter the partial volume effect.
Please note that while it is understood in the art to which this invention relates that nuclear medicine imaging refers to imaging using radioactive tracers, typically introduced into the body of the subject being imaged, the applicant wishes to emphasize that the invention also applies to radiation-based medical imaging, where the rays may include other forms of radiation in medical imaging, for example, X-rays produced in an X-ray tube outside the body being imaged, or a radioactive source outside the body producing gamma rays for imaging the body. Images can be produced from ingested radioactive tracers or from a source outside the body being imaged.